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8/3/2019 John A. Stella et al- Tissue-to-Cellular Level Deformation Coupling in Cell-Microintegrated Elastomeric Scaffolds
1/18
TISSUE-TO-CELLULAR LEVEL DEFORMATION COUPLING IN
CELL-MICROINTEGRATED ELASTOMERIC SCAFFOLDS
John A. Stella1, Jun Liao1,*, Yi Hong2,3, W. David Merryman1,**, William R. Wagner1,2,3, andMichael S. Sacks1,2
1Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA United States
2McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA United States
3Departments of Surgery and Chemical Engineering, University of Pittsburgh, Pittsburgh, PA United States
Abstract
In engineered tissues we are challenged to reproduce extracellular matrix and cellular deformation
coupling that occurs within native tissues, which is a meso-micro scale phenomenon that profoundlyaffects tissue growth and remodeling. With our ability to electrospin polymer fiber scaffolds while
simultaneously electrospraying viable cells, we are provided with a unique platform to investigate
cellular deformations within a three dimensional elastomeric fibrous scaffold. Scaffold specimens
micro-integrated with vascular smooth muscle cells were subjected to controlled biaxial stretch with
3D cellular deformations and local fiber micro-architecture simultaneously quantified. We
demonstrated that the local fiber geometry followed an affine behavior, so that it could be predicted
by macro scaffold deformations. However, local cellular deformations depended non-linearly on
changes in fiber microarchitecture and ceased at large strains where the scaffold fibers completely
straightened. Thus, local scaffold microstructural changes induced by macro-level applied strain
dominated cellular deformations, so that monotonic increases in scaffold strain do not necessitate
similar levels of cellular deformation. This result has fundamental implications when attempting to
elucidate the events of de-novo tissue development and remodeling in engineered tissues, which are
thought to depend substantially on cellular deformations.
INTRODUCTION
Tissue engineering aims to recapitulate native tissue structure, composition, and mechanical
function in a controlled and reproducible manner to overcome the limitations of current medical
therapies. The successful development of tissue engineered therapies rests in large part on our
ability to employ new materials, manufacturing and processing techniques, and control of
cellular mechanobiology. These are all predicated on a strong fundamental knowledge of native
cellular and scaffold functional coupling. Cells perceive and react to their mechanical
environment through adhesion to local substrates as well as cellular deformations induced by
the surrounding mechanical environment. These complex interactions likely play a critical role
in the mechanotransduction of proteins required for cell viability and proliferation, and have
For correspondence: Michael S. Sacks, Ph.D., 100 Technology Drive, Room 234, University of Pittsburgh, Pittsburgh, PA 15219, Tel:412-235-5146, Fax: 412-235-5160, email: [email protected].*Current address: Department of Agricultural and Biological Engineering, Mississippi State University**Current address: Department of Biomedical Engineering, University of Alabama at Birmingham
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NIH Public AccessAuthor ManuscriptBiomaterials. Author manuscript; available in PMC 2009 August 1.
Published in final edited form as:
Biomaterials. 2008 August ; 29(22): 32283236. doi:10.1016/j.biomaterials.2008.04.029.
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direct implications towards the development of engineered tissues. However, contemporary
knowledge is lacking as to how externally applied traction forces at the tissue level are
transmitted to the cell in both native and engineered tissue systems. This is in part due to
technical difficulties with performing studies in native tissues, which can be confounded by
complex hierarchical structures, vasculature, and multiple cell types.
In engineered tissues, various technologies [14] have been used to construct layered tissues
but have limited abilities to control cellular orientation, while other methods have generated3D pore structures that show promise in modulating cellular function [5,6]. Yet, irrespective
of the particular approach, a key issue remains on how scaffold-level structures and
deformations translate to cell level deformations. Cellular deformations are known to strongly
affect cellular biosynthetic responses in various tissues [7], yet this important aspect remains
largely unexplored in the tissue engineering literature. In native tissues, cells and extracellular
matrices (ECM) are regularly arranged to form hierarchical architectures that span multiple
length scales to form complex three dimensional tissue systems. Clearly, operational scales
have evolved to match the natural scales, and one should focus on the appropriate scales most
relevant for the problem at hand [8].
For engineered tissue applications, it is our aim to address the relation between deformations
that occur at the cellular level and those that occur at the tissue level due to organ-level forces
in a synthetic fibrous scaffold densely integrated with cells. Herein, we utilized poly (esterurethane) urea (PEUU) as the elastomeric scaffold material due to its biodegradable and
cytocompatible properties [9]. With our ability to incorporate viable cells distributed
throughout the scaffold via concurrent electrospraying and electrospinning (ES) of PEUU fiber
scaffolds (Fig. 1-a), we are provided a unique platform to investigate the 3D deformation
response of cells seeded in a controlled fiber architecture in-situ [9,10]. Moreover, electrospun
scaffolds have structural features that span from m to multi-mm scales, and are thus analogous
to native soft tissue structures. We hypothesize that cellular deformations within ES-PEUU
scaffolds are likely a complex function of scaffold mechanical properties, architecture, and
cellular coupling with the surrounding polymer fibers. Our objective was therefore to quantify
cell deformations and relate them to changes in specific scaffold structures in response to
externally applied large (finite) tissue-level strains utilizing real-time laser scanning confocal
microscopy (LSCM).
METHODS
Specimen fabrication and preparation
For a detailed account of the concurrent electrospraying and electrospinning process, the
interested reader is referred to the previous work of Stankus et al.[911]. Briefly, the
electrospinning process produces continuous fiber scaffolds exhibiting a wide range of
mechanical properties while providing a suitable environment for cell proliferation and growth
[9,1216]. Electrospinning involves the deposition of a solubilized polymer across a large
voltage gradient onto a collection surface, with the polymer solution delivered via a small
capillary tube where the applied electric potential induces a charge accumulation on the surface
of the expelled polymer (Fig. 1-a). Once electrostatic forces exceed viscoelastic forces within
the polymer solution and surface tension, a fine jet is ejected towards the collection surface
while being accompanied by rapid solvent evaporation. The result is a mat of continuous
polymer fibers that, depending upon the processing parameters and polymer utilized, can
exhibit mechanical behavior very similar to native soft collagenous tissues [15,17].
For the current study, electrospraying of rat VSMCs was accomplished by feeding 1107 cells/
ml into a sterilized capillary charged at 7 kV and located 4 cm from the target mandrel (4.7
mm diameter) concurrent with PEUU/1,1,1,3,3,3-hexafluoroisopropanol solution (12 wt%)
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deposition from a capillary charged at 12 kV and located 20 cm from the target mandrel (Fig.
1-a). The mandrel was charged at 4 kV and was rotated at 150 rpm while translating 8 cm
along the z-axis at 0.15 cm/s. After 30 min of electrospinning and electrospraying, the
microintegrated tube was removed from the mandrel and cultured in a spinner flask at a rotation
rate of 15 rpm for 2448 h prior to imaging. With the capability of simultaneously
distinguishing living VSMC nuclei stained with DRAQ5 and PEUU fibers below the specimen
surface, an inverted Laser Scanning Confocal Microscope (LSCM, Olympus Fluoview 1000)
was chosen.
Biaxial stretch device and testing
Biaxial modes of deformation are well known to better approximate physiological
deformations of planar anisotropic tissues [18]. Moreover, we were interested how the scaffold
fibers stretch under biaxial deformations without the confounding effects of simultaneous fiber
rotations [19]. A biaxial stretcher was thus custom designed for use with an inverted laser
scanning confocal microscope (LSCM, Olympus Fluoview 1000). The biaxial stretcher
consisted of a polycarbonate specimen chamber with two pairs of orthogonally positioned lead
screws (Fig. 1-b). Two loops of 5-0 polyester suture of equal length were attached from the
lead screws to each side of the specimens with four custom made stainless steel hooks.
Subsurface specimen imaging was facilitated by a glass cover slip (Electron Microscopy
Sciences Hatfield, PA, USA) mounted to the stretching device spanning a 30 mm diameter
opening in the chamber beneath the specimen. This system was capable of applying both strip
biaxial and equibiaxial stretch regimes. Strip biaxial stretch refers to a special loading regime
wherein the stretch along one axis is increased while the orthogonal axis is constrained such
that no deformation occurs. Conversely, equibiaxial deformation is defined as an equal level
of stretch in the two orthogonal specimen directions.
Confocal imaging
An inverted Laser Scanning Confocal Microscope (LSCM) was chosen to observe living cells
and scaffold in situ, with images taken via a coverslip window below the specimen. Just prior
to imaging, specimens measuring 12 mm square were stained with DRAQ5 for 30 minutes, a
far red-fluorescing DNA probe for viable cells. DRAQ5 has been demonstrated not to be toxic
to cells with nuclear DNA staining being stable up to 4 h [20]. Moreover, DRAQ5 proved to
be DNA staining specific in our experiments allowing us to image live cells concurrently withthe auto-fluorescing PEUU fibers. Cell nuclear images were taken under CY5 channel and the
scaffold was imaged under CY3 channel via polymer autofluorescence. For each strain level,
randomly selected regions were imaged at a subsurface depth ranging from approximately 15
75 m. For three specimens, 2586 image stacks were obtained with section thicknesses of
0.972 m 2.831 m with image intervals ranging from 0.6 m 0.8 m, producing a net
thickness of 17.4 to 68.0 m. This imaging approach produced good quality representations
of both the nuclei and fibers under stretch (Fig. 1-c,d).
Cellular deformations and scaffold stretch protocols
Nuclear aspect ratio (NAR) has been used as a measure of overall cellular deformation in native
tissues (see discussion) [2124]. For example, aortic heart valve tissues are populated with
valvular interstitial myofibroblast cells (AVIC) which serve to maintain the valve leafletextracellular matrix (ECM) [25, 26]. We have shown [22] that AVICs undergo very large
deformations with changes in NAR from 1.8 in the unloaded state to 5.0 under full transvalvular
pressure, which is likely related to their biosynthetic function (Fig. 2-a). Thus, to quantify the
3D geometry of the living micro-integrated cell nuclei within the scaffolds, complete LSCM
image stacks spanning 1575 m were obtained for specimens in the reference and deformed
configurations (n=3). From the image stacks we reconstructed the complete 3D VSMC nuclear
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surface and observed the cell nuclei consistently exhibited a scalene ellipsoidal geometry (Fig.
2-b). Amira (Mercury Computer Systems, Inc., Carlsbad, CA, USA) was used to stack the
registered images and segment cell nuclei manually. Next, MATLAB (The MathWorks, Inc.,
Natick, MA, USA) was employed to perform principle component analysis (PCA) on the
segmented cell nuclei. The full 3D nuclear orientation was computed using principle
component analysis of the reconstructed nuclear surface by determining the angles of the three
principle nuclear directions with respect to the specimen coordinate axes (, , and
respectively (Fig. 2-c). Cell nuclei aspect ratios of 2D images were quantified using thecommercial software Sigma Scan Pro (SPSS Inc., Chicago, IL, USA). For each strain state,
the NAR for 80 cells were analyzed with data being presented as AVG SEM.
As in our previous work with soft tissues [19,27], we utilized equi-biaxial deformation states
(i.e. were the specimen is loaded such that the axial strains, PD and XD, remain equal), to
allow extensional strains without fiber rotations. This was accomplished with both the scaffold
fibers imaged by LSCM and also by scanning electron microscopy (SEM) to provide additional
detail. Next, in order to induce the largest changes in fiber stretch and rotation, as well as
cellular deformations, strip biaxial strain states were chosen wherein extensional strain is
applied in one direction while the strain the other direction is held at zero. Specific deformation
levels utilized included the unloaded configuration followed by states of increasing strip biaxial
strain ranging from 5% to 102%, and equi-biaxial strain states of 10% to 50% in 10%
increments (n=15 total).
TEM cellular imaging
To verify that NAR could be used as an index of cellular deformations using micro-integrated
VSMCs, transmission electron microscopy (TEM) was used to compare and quantify the
geometries of the VSMC membrane and nucleus. Due to the polymeric nature of the cell
integrated scaffolds it was not possible to fix specimens in their deformed states. As a result,
it was only possible to image cells in a free floating configuration. Cell-scaffold construct
sections were cut parallel to the preferred and cross-preferred fiber directions for imaging. In
total, 22 cells were imaged in the preferred direction while 10 were imaged in the cross-
preferred direction. The major and minor axis lengths were then measured manually with the
image analysis software Sigma Scan Pro (Systat Software, Inc., San Jose, CA USA). NAR and
cell membrane aspect ratios of cells in unstrained PEUU scaffolds were not observed to be
significantly different in either the PD or XD directions (pPD = 0.745 and pXD = 0.213).
Scaffold fiber kinematical analysis
Fiber orientation was also quantified from the LSCM images (Fig. 1-c,d) using custom image
analysis software in the unstrained and deformed configurations. Methods for this custom
image analysis software have been previously presented in detail [17, 28, 29]. Essentially, the
software directly produces statistical distributions of fiber orientation probability, R(), from
all measured fibers over the range of all possible orientations, 9090. Next, to gain a
better understanding of how the ES-PEUU fibers deform as a function of macroscopic scaffold
strain, we utilized an affine fiber deformation transformation model to evaluate if the ES-PEUU
followed this model, as we have done for collagenous scaffolds [19]. In brief, the measured
in-plane deformation gradient tensor (F) was first computed from four graphite markers
adhered to the surface of the specimen (Fig. 1-b). Polar decomposition ofF was then performedto remove small rigid body rotations that occur in real time using
(1)
where U and R are the right stretch and rotation tensors, respectively [30]. Essentially, eqn. 1
decomposes the net deformation into rigid body rotations and stretch contributions. In the
current study we found that rigid body rotations were negligible, so that FU. Thus, assuming
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a homogeneous, affine deformation the rotation of fibers originally at an angle state to the
angle in the deformed state is given by [19]
(2)
where FPD, FXD, and FPD-XD are the extensional and shear stretches along the preferred and
cross-preferred directions, respectively (Fig. 1-b). If the assumption of affine deformation
holds, the angular fiber distribution in the deformed state, R(), can be related to its originalangular distribution R() using
(3)
RESULTS
Fiber kinematics
Comparison of both the LSCM and SEM imaging modes demonstrated the LSCM provided
comparable levels of detail, and at least sufficient fiber density to obtain accurate orientation
information (Fig. 3-a,b). Upon stretch, a substantial increase in fiber local straightening was
observed with increased equi-biaxial stretch (Fig. 3-c,d). Interestingly, at high strain levels
there was an upper bound of fiber alignment observed, corresponding to a transition from anensemble of tortuous fibers seen in the non-deformed specimen to a web-like architecture as
manifested by a scaffold with straight, interconnected fibers. Thus, ES-PEUU fibers do not
completely loose there tortuosity with strain, and appear to be bonded together. Next, fiber
orientation distributions indicated that when unloaded, the fiber network returned immediately
back to its original orientation (Fig. 4-a), clearly indicating the elastic nature of the scaffold.
Moreover, results for the equi-biaxial tests confirmed that no measurable fiber realignment
occurred, as what should occur if the fibers follow the affine model (Fig. 4-b).
In further exploration of the changes in fiber orientation in the strip biaxial strain tests, the fiber
orientation distribution R() was observed to become more random with increasing strain (Fig.
5-a). This result was consistent with the fiber network deforming orthogonally from its
preferred direction. Next, we assessed whether the ES-PEUU fibers deform according to an
affine rule (eqn. 1eqn. 3) in non-equibiaxial strain states. Interestingly, at both low (Fig. 4-b)and high (Fig. 4-c) strain levels the affine rule predicted good agreement with the experimental
data. Thus, due to their high inter-connectivity (e.g. Fig. 3) the fibers in an average sense
deform in accordance to the bulk strain field.
3D cellular deformations and the relation to local fiber kinematics
From the 3D cellular nuclei reconstructions, we noted that for all measured cells (n = 20) the
angle , which defined the out-of-plane orientation of the nucleus (Fig. 2-c), was 5.92 1.57
in the unloaded configuration. This indicated that the micro-integrated cells were aligned
mainly in the plane of the scaffold. Moreover, under increasing strip biaxial deformation the
major principle direction rotated in the direction of deformation (Fig. 6), suggesting substantial
mobility with strain.
Next, we examined the interrelationship between the micro-integrated cellular deformations
and the changes in scaffolds fibrous structure. This was facilitated by simplifying imaging and
analysis to use single section imaging sets to allow for multiple states of deformation to be
obtained with same specimen while maintaining cellular viability, which did not result in
measureable information loss. No dependency was observed in the NAR-strain relationship
for changes in the sequence of deformation, indicating that the cellular-scaffold deformations
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were largely elastic in nature. We noted in particular that, under strip biaxial stretch, cells
micro-integrated into ES-PEUU scaffolds exhibited a rapid increase in NAR until
approximately 60% strain (Fig. 7-a). At strain levels higher than 60%, a plateau in NAR was
observed (Fig. 7-a). Moreover, this behavior did not depend upon the direction of strip biaxial
stretch (PD vs. XD stretch, Fig. 7-a). Interestingly, the cessation in NAR change with strain
corresponded closely with complete straightening of the local fibrous structure (Fig. 3). This
result suggested local fiber straightening, as opposed to macro-tissue level deformations, was
the dominant mechanism for inducing cellular deformations. It should be also noted that thisbehavior is in stark contrast to that observed in the native tissues, wherein the largest changes
occurred at higher stress levels (Fig. 7-b).
DISCUSSION
Choice of scaffold in the study of cellular deformations in engineered tissues
Electrospinning produces continuous fiber scaffolds exhibiting a wide range of mechanical
properties, while also providing suitable surfaces for cell proliferation and growth [9,1216].
The electrospinning process produces scaffold sheets of PEUU fibers (0.8 m in diameter)
approximately 300400 m thick to approximate the scale and mechanical behavior of native
extracellular matrix, including the ability to undergo large deformations and fully recover when
unloaded (i.e. nearly elastic behavior). While electrospinning can fabricate scaffolds that
possess ECM-like structures, this morphology also results in pore sizes that are generallysmaller (
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the large amount of fiber straightening for pressures below ~1 mmHg, followed by substantial
increases in AVIC NAR from 4 to 90 mmHg [22] (Fig. 6-b). Taken as a whole, cell responses
to tissue level stresses are modulated through complex micromechanical and fiber-compaction
effects that occur under physiological stress levels.
Role of fiber micro-mechanics in cell deformations
The cell micro-integrated ES-PEUU scaffolds exhibited micro-fiber morphologies and
kinematics that were shown to directly influence local cellular deformations. For instance, inthe unstrained configuration the ES-PEUU fibers exhibited a tortuous architecture which
transitioned to a web-like network of straight, interconnected fibers at high levels of strain.
The deformations of the micro-integrated VSMC were found to be primarily mediated by this
phenomenon. The VSMC integrated ES-PEUU constructs underwent fully recoverable large
deformations akin to many native tissues. Moreover, while a non-linear relation between the
tissue strain and NAR was observed for both the aortic valve and cell integrated ES PEUU
(Fig. 7), the underlying micro-mechanical mechanisms were clearly different. Initially, the
integrated VSMCs exhibited a rapid increase in NAR as fibers straightened and tortuosity was
reduced. Once the PEUU fibers became straightened and the architecture transitioned to an
interconnected web like structure, changes in NAR were observed to plateau. In contrast to the
compression mediated deformations observed in the AVIC (Fig. 7-b), the microintegrated
VSMC deformation was mediated by the local reduction of tortuosity or straightening of the
PEUU fibers (Fig. 7-a). Thus, cell-scaffold interactions can be subtle and can bring about
significantly different deformation behaviors.
Broader considerations
Despite its early successes, tissue engineering continues to face challenges in repairing or
replacing tissues that serve a predominantly biomechanical function. An evolving discipline
called functional tissue engineering seeks to address the development of load-bearing
structures in several ways [35]. In particular, a critical subset of native tissue mechanical
properties must be selected and prioritized as design objectives. This subset is important, given
that the mechanical properties of the designs are not expected to completely duplicate the
properties of the native tissues. Increasing evidence suggests that mechanical stress, as well as
other physical factors, may significantly increase the biosynthetic activity of cells in
bioartificial matrices. Clearly, the effects of physical factors on cellular activity must bedetermined in engineered tissues. Knowledge of these signals may shorten the iterations
required to replace a tissue successfully and direct cellular activity and phenotype toward a
desired end goal. Ultimately, incorporating each of these principles of functional tissue
engineering should result in safer and more efficacious repairs and replacements.
Substantial difficulties are imposed in functional tissue engineering since there are multiple
length scales with complex architectures, modes of deformation, and biochemical stimuli
which work synergistically to determine physiologic responses. It is generally accepted that
both chemical and mechanical factors modulate cell biosynthesis when producing extracellular
matrix [34,3638]. Yet, the exact microstructural characteristics of the scaffolds will have a
profound influence on cellular function. For example, collagen and fibrin gels are a popular
choice in engineered scaffolds [3943]. Yet, while possessing desirable characteristics such as
cyto-compatibility and anisotropic mechanical properties, their tensile properties have not yetachieved that of native tissues. Moreover, native fibrous protein based gels are largely
composed of short-range fibers, whereas native connective tissues are primarily composed of
long fibers, which can span multi-mm scales and simultaneously facilitate flexibility and high
tensile strength.
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As an alternative, biodegradable synthetic polymers as scaffolds to support and encourage
tissue regeneration have also proven successful. For example, Freed et al. employed non-woven
meshes of polyglycolic acid as a scaffold for seeded chondrocytes and successfully showed
the ability to regenerate cartilaginous tissues composed of glycosaminoglycans and collagen
[44]. As mentioned previously, electrospinning is a versatile process and slight alterations in
the manufacturing process enable the production of scaffolds with a wide array of fiber
morphologies (i.e. fiber diameter, porosity, packing density, orientation, etc) which directly
influence bulk mechanical properties [45,46]. Controlled mechanical anisotropy, for example,is attained by using a rotating collection surface which induces a preferred fiber direction as
the rotational speed of the collector increases [17]. This ability is extremely beneficial in
mimicking native tissue architecture and has even been shown to approximate the highly
nonlinear biaxial mechanical response of collagenous soft tissues, such as the native porcine
pulmonary valve leaflet [17].
In light of these studies, the importance of scaffold-cell mechanical coupling becomes readily
apparent in the rational design of engineered tissues. Moreover, the unique micromechanics
of various scaffolds induce different cell deformation responses, which could correlate to
substantial changes in cell proliferation and function. This phenomenon is illustrated by three
grey bars placed at intervals of 20% strain which each correspond to unique NAR values.
Instead, the change in cell deformation (or NAR)-strain relationship explored here can be used
to guide future engineered tissue applications for these integrated ES PEUU scaffolds.
Limitations
It should be noted that we investigated the response cell integrated ES PEUU scaffolds cultured
for 48 hours, which contain minimal functional ECM. The forming ECM within the scaffold
and its subsequent micromechanical effects will likely have a substantial effect on cellular
deformation, and could prove critical in defining the long term behavior of cell function within
integrated electrospun scaffolds. In addition, much of the complex interactions between VSMC
and polymeric scaffolds remain unknown, including the nature of cellular focal adhesions to
the PEUU matrix. It is likely that the nature and extent of adhesion sites between a cell and its
surrounding matrix will directly impact manifestations of biosynthetic activity.
Conclusions
It remains a critical goal to understand how organ-level deformations, such as cyclic stretch,
bending, and shear [47,48], translate to microstructural deformations and ultimately as cellular
mechanical stimuli. A great deal can be learned about the mechanical modulation of functional
tissue from ES-PEUU scaffolds, since they capture some aspects of native tissue microstructure
and exhibit the ability to endure large deformations while recovering completely. Moreover,
these issues become more challenging in the formation of actively contracting tissues, where
cellular contraction must be coordinated at higher scales in order to provide large stresses
required for organ contraction. The emergence of new materials and processing methods will
clearly be required to meet these functional needs.
ACKNOWLEDGMENTS
Funding for this work was provided by NIH R01s HL68816 and HL69368. John Stella was partially supported by theNIH-NHLBI training grant (T32-HL76124) entitled "Cardiovascular Bioengineering Training Program." Additional
support for Drs. Jun Liao and W. David Merryman came from American Heart Association Grant-in-Aid (0565346U)
and Pre-doctoral Fellowship (0515416U), respectively.
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Figure 1.
Fabrication and imaging under biaxial stretch of cell integrated elastomeric scaffolds. (a) Cell-
scaffold constructs were fabricated via concurrently electrospinning polymer andelectrospraying cells onto a 4.7 mm diameter rotating mandrel. (b) Biaxial stretch device used
to deform cell micro-integrated ES-PEUU scaffolds. (c,d) As the scaffold construct underwent
strip biaxial deformation, the cell nuclei (blue) were observed to elongate and the fibers become
straightened (red).
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Figure 2.
Nuclear aspect ratio deformation and orientation. (a) In native tissues such as the aortic valve,
cell nuclei undergo very large changes in aspect ratio induced by tissue stretch and compaction.Here aortic valve tissue cross-sections are shown in the unloaded (0 mmHg) and full
physiological loaded (90 mmHg) states. (bc) 3D reconstructions of the micro-integrated
vascular smooth muscle cells in the unloaded configuration, whose 3D orientation was
quantified with respect to the specimen PD and XD axes by the angles ,, and .
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Figure 3.
Changes in ES-PEUU fiber microarchitecture under biaxial stretch. When strained equally in
both orthogonal directions, PEUU fibers were observed by (a,c) SEM and (b,d) laser scanning
confocal microscope to transition from a tortuous configuration in the unstrained state to an
interconnected web-like architecture at high strains state.
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Figure 4.
(a) When strained equi-biaxially and unloaded, the ES-PEUU fibers were observed to return
to their original orientations, suggesting an elastic behavior. Moreover, when strained equi-
biaxially by at various levels, not detectable change in orientation occurred, as predicted by
affine fiber kinematics.
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Figure 5.
ES-PEUU fiber architecture responds to strain in an affine sense. To induce the greatest fiber
rotations, we strained the scaffolds along the XD direction to a mid-level (1525%) and high(>65%), and noted a gradual shift of the fiber orientation distribution function R() to a random-
like distribution (where R() = 1/) (a). When we compared the R() to that predicted by an
affine model (see Methods), good agreement was found at both (b) mid-level and (c) high level
strains. This suggests that ES-PEUU fiber deform in a bulk sense with the macro-scaffold
strains.
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Figure 6.
Micro-integrated cell rotations during deformation. As the cell-microintegrated scaffolds were
deformed under strip biaxial stretch, the integrated cells principal component directions wereobserved to rotate and become more uniformly aligned in the direction of deformation. Here,
this phenomenon is shown with all measured major principal directions shown as unit vectors.
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Figure 7.
NAR change with deformation is closely related to local fiber microarchitecture. (a) A
composite of all NAR measurements (mean s.e.m) demonstrated a rapid increase to ~60%strain, after which a plateau was observed with further strain increases, indicating that nuclei
deformations are dominated by local fiber straightening. (b) Cells in native tissues, in contrast,
exhibit a very different response when exposed to physiologically stress levels. In this example,
when compared to local collagen fiber alignment AVIC nuclear deformations undergo
relatively little change at low transvalvular pressures (i.e. low tissue stresses), whereas at high
transvalvular pressures (i.e. high tissue stresses) they are dominated by fiber compaction
effects. Modified from [22].
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